Technical Field
The present invention relates to a device for delivery of material or stimulus to targets within a body to produce a desired response, and in particular to a device including a number of projections for penetrating a body surface. The invention can also relate to devices for delivering bioactive substances and other stimuli to living cells, to methods of manufacture of the device and to various uses of the device.
Description of the Related Art
The reference in this specification to any prior publication (or information derived from it), or to any matter which is known, is not, and should not be taken as an acknowledgment or admission or any form of suggestion that the prior publication (or information derived from it) or known matter forms part of the common general knowledge in the field of endeavor to which this specification relates.
In recent years, attempts have been made to devise new methods of delivering drugs and other bioactive materials, for vaccination and other purposes, which provide alternatives that are more convenient and/or enhanced in performance to the customary routes of administration such as intramuscular and intradermal injection. Limitations of intradermal injection include: cross-contamination through needle-stick injuries in health workers; injection phobia from a needle and syringe; and most importantly, as a result of its comparatively large scale and method of administration, the needle and syringe cannot target key cells in the outer skin layers (FIG. 11(a)). This is a serious limitation to many existing and emerging strategies for the prevention, treatment and monitoring of a range of untreatable diseases.
The skin structure is shown in FIG. 11, with a summary of key existing delivery methods. Non-invasive methods of delivery through the skin have been used, including patches, liquid solutions and creams. Their success is dependent upon the ability to breach the semi-permeable stratum corneum (SC) into the viable epidermis. Typically, larger biomolecules are unable to breach this barrier.
Alternatively, there are many more “invasive” means to breach the SC for pharmaceutical delivery to the viable epidermis. Simple methods include: tape stripping with an abrasive tape to or sandpaper and the application of depilatory agents. Amongst the more advanced technologies are electroporation, ablation by laser or heat, radiofrequency high voltage currents, iontopheresis, liposomes, sonophoresis. Many of these approaches remain untested for complex entities such as vaccines and immunotherapies. Moreover, they do not specifically deliver entities within key skin cells.
Needle-free injection approaches include the high-speed liquid jet injector, which had a rise and fall in popularity in the mid twentieth century—and has recently seen a resurgence (Furth, P. A., Shamay, A. & Henninghausen, L. (1995) Gene transfer into mammalian cells by jet injection. Hybridoma, 14:149-152.). However, this method delivers jets of liquid to the epidermis and dermis (labelled (c) in Fig A), usually with a diameter >100 μm and not within key cells. Furthermore, as a result of the concentrated jet momentum, many skin cells die. Delivery into the dermis also leads to patients reporting pain from injection.
The ballistic, needle-free delivery of microparticles (or gene gun) offers a route for delivering biological agents directly into cells of the skin. In this needle-free technique, pharmaceutical or immunomodulatory agents, formulated as or coated to particles, are accelerated in a high-speed gas jet at sufficient momentum to penetrate the skin (or mucosal) layer and to achieve a pharmacological effect. A schematic of microparticles in the skin following ballistic delivery is shown in FIG. 11(b). The ability of this “scatter gun” approach to deliver genes and drugs to epidermal cells is highly limited and sensitive to biological variability in skin properties on the dynamic high strain rate ballistics process. These effects are discussed in Kendall, M. A. F., Rishworth, S., Carter, F. V. & Mitchell, T. J. (2004) “The effects of relative humidity and ambient temperature on the ballistic delivery of micro-particles into excised porcine skin.” J. Investigative Dermatology, 122(3):739-746.); and Kendall, M. A. F., Mitchell, T. J. & Wrighton-Smith, P. (2004) Intradermal ballistic delivery of microparticles into excised human skin for drug and vaccine applications. J. Biomechanics, 37(11):1733-1741.
First, the ballistic delivery of particles into the skin to target epidermal cells is extremely sensitive to the small variations in the stratum corneum-including the stratum corneum thickness, which varies massively with body site, age, sex, race and exposure to climatic conditions. (The quasi-static loading of skin with micro-nanostructures would be less sensitive to these differences).
Second, it has been shown that even when all these parameters are strictly controlled—and the only parameter varied is the climatic relative humidity (15%-95%), or, independently, temperature (20° C.-40° C.)—the result is a large variation in penetration depth. These results are shown in FIG. 12, with particle penetration as a function of ambient relative humidity (FIG. 12(a)) and ambient temperature (FIG. 12(b)) plotted along with theoretical calculations of particle penetration and measured stratum corneum thickness. This variation alone is significant and sufficient to make the difference between particles breaching the stratum corneum, or not.
The compound effect of these two (and other) sources of variability is the gene gun/biolistics process does not consistently target epidermal cells-leading to inconsistent biological responses (e.g., in DNA vaccination).
Interestingly, it has also been shown that the high strain-rate loading of the skin under ballistic particle impact (approximately 106 per second) increases the stratum corneum breaking stress by up to a factor of 10 compared to quasi-static values-due to a ductile-to-brittle change in the skin mechanical properties. This means that the tissue is more difficult to penetrate as the particle impact velocity is increased. Therefore it is desirable to devise a way to deliver micro/nanostructures to the skin at lower strain-rates than the ballistic approach to exploit the weaker stratum corneum.
When the microparticles are delivered to the skin, it is unclear whether there are any adverse longer term effects. For example, in the case of insoluble particles, many of them slough off with the usual skin turnover. However, gold particles have been detected in the lymph nodes following ballistic particle delivery-presumably by migration with Langerhans cells. Uncertainty of adverse effects of these delivered materials would be removed by delivery routes that do not leave such materials in tissue site.
Moreover, when the microparticles successfully target cells, there is a significant probability they kill the cells they target. Consider a typical ballistic delivery condition: over 1 million 2-3 μm diameter gold particles coated in DNA to the skin at 400-600 m/s, over a target diameter of 4 mm (Kendall, M. A. F., Mulholland, W. J Tirlapur, U. K., Arbuthnott, E. S. & Armitage, M. (2003) Targeted delivery of micro-particles to epithelial cells for immunotherapy and vaccines: an experimental and probabilistic study. 6th International Conference on Cellular Engineering. Aug. 20-22, 2003, Sydney, Australia.). Reported experiments with these conditions using cell death stains (ethidium bromide/acridin orange) show that microparticles impacting the skin do kill cells (McSloy, N. J Raju, P. A. & Kendall, M. A. F. (2004) The effects of shock waves and particle penetration in skin on cell viability following gene gun delivery. British Society for Gene Therapy, 1st Annual Conference. Oxford, UK, Mar. 28-30, 2004; Raju, P. A. & Kendall, M. A. F. (2004) Epidermal cell viability following the ballistic delivery of DNA vaccine microparticles. DNA Vaccines 2004—the Gene Vaccine Conference. 17-19 Nov. 2004, Monte Carlo, Monaco.).
FIG. 13A shows the percentage of cells that had membrane rupture (i.e., death) as a function of the localized particle channel density. In FIG. 13B we see schematically the way the data in FIG. 13A was achieved, relating “tracks” left by particle penetration to the death of cells in a layer of the viable epidermis. Clearly, FIG. 13A shows that at a channel density above 0.01 channels/micron, all the cells in that layer are dead. Indeed, FIG. 13C shows that cells are killed when the particle is passing up to 10 μm outside of the cell boundary. The mechanism of cell death is due to the propagation of stress and shock waves in the skin generated by the rapid deceleration of the microparticles (McSloy et al. (2004)). The rapid rise time of these stress waves in the skin, and their magnitude both contribute to cell death and the results are consistent with the findings reported by Doukas, A. G. & Flotte, T. J. (1996). Physical characterization and biological effects of laser-induced stress waves. Ultrasound in Medicine and Biology, 22(2):151-164. The effects of shock waves and particle penetration in the skin on cell viability following gene gun delivery. Masters Thesis, Department of Engineering Science, University of Oxford.). This mechanism of ballistic particle penetration killing cells negatively affects the ability of the direct and efficient delivery of genes and drugs to the cells.
This cell death effect of ballistic particle delivery could be reduced by significantly decreasing the particle size to the nanometer regime-thereby reducing the stresses on the cells. However, another limitation of the gene gun is that it is unsuitable in delivering sub-micron sized particles to cells. This is illustrated by the following. As reported (Kendall, M. A. F., Mitchell, T. J. & Wrighton-Smith, P. (2004) Intradermal ballistic delivery of micro-particles into excised human skin for drug and vaccine applications. J Biomechanics, 37(11):1733-1741), and shown in FIG. 14, ballistic particle penetration is proportional to the particle impact parameter, pvr, which is the product of the particle density (ρ), velocity (v) and radius (r). This parameter is also proportional to the particle momentum per-unit-area, which has been shown to drive the mechanism of particle penetration depth (Mitchell et al. (2003)). From FIG. 14, we see a 1 μm radius gold particle (density 18000 kg/m3) would need to impact the skin at ˜600 m/s in order to penetrate to reach cells ˜20 μm into the skin.
Experimental results show that reducing the particle radius, say, by an order of magnitude, to 100 nm, and placing it in a standard biolistic device leads to negligible particle impact in the skin. Indeed from FIG. 14 we see delivery to a 20 μm depth would need an impact velocity of ˜6000 m/s, which is impractical for two reasons: 1) these hypervelocity conditions can not be safely achieved with a system configured for human use (they are usually achieved with massive free-piston shock tunnels); 2) even if 6000 m/s was obtained in the free-jet, a gas impingement region above the skin would seriously decrease the particle velocity—it is possible that the particle would not even hit the skin at all. Interestingly if a method was conceived to safely and practically deliver nanoparticles to the skin at higher velocity (e.g., the stated case of an 100 nm radius gold particle at a velocity of ˜6000 m/s), the cell death benefit of smaller scale would be offset by higher peak stresses-killing more cells—and higher strain rates that are likely to further “toughen” the skin, making delivery even more difficult.
In conclusion, these collective facts rule out the gene gun as a viable option for delivering nanoparticles and therefore precludes it from many of the developments in biomolecules, drugs and sensors at this scale.
The huge research effort in micro- and nanotechnologies provides tremendous potential for simple and practical cell targeting strategies to overcome many limitations of current biolistic (and other) cell targeting approaches. For example, FIG. 11(c) shows that the most conceptually simple and appealing approach to gene delivery is the direct injection of naked DNA to live cell nuclei at a sub-micrometer scale that does not adversely damage the cell (Luo, D. & Saltzman, W. M. (2000) Synthetic DNA delivery systems. Nature Biotechnology, 18:33-36). Cell death is minimized by both the sub-micrometer scale of the injector and the low, quasi-static strain-rate of the probe (compared to ballistic delivery) resulting in low stress distributions. Although this is a very efficient gene and bioagent delivery route, the to drawback is that such precise targeting by direct microinjection can only be achieved one cell at a time and with great difficulty to the operator in vivo. Hence, the method is slow, laborious and impractical.
Researchers have overcome some of these disadvantages for transdermal drug delivery by fabricating arrays of micrometer-scale projections (thousands on a patch) to breach the stratum corneum for the intradermal delivery of antigens and adjuvants to humans and other mammals.
In the scientific literature, the first description of this technique appears to be the paper Microfabricated Microneedles: A Novel Approach to Transdermal Drug Delivery. S. Henry et al, J. Pharmaceutical Sci. vol. 87(8) p 922-925 (1998), with the accompanying patent of U.S. Pat. No. 6,503,231. The objective of U.S. Pat. No. 6,503,231 is to provide a microneedle array device for relatively painless, controlled, safe, convenient transdermal delivery of a variety of drugs and for biosampling. This is achieved by the microneedles breaching the tissue barrier (e.g., for skin: the stratum corneum) and then the therapeutic or diagnostic material is injected through the hollow microneedles into the tissue. Specifically, in claim 1 of U.S. Pat. No. 6,503,231, it is stated that the microneedles are to be hollow, with a length of 100 μm-1 mm, and claim 3 states the width of 1 μm-100 μm, with subsequent claims stating ways the hollow needles can be connected to reservoirs for the injection of liquids, fabrication methods, materials and examples of drugs to be delivered. Thus, U.S. Pat. No. 6,503,231 describes a patch suitable for delivering materials and/or energy across tissue barriers. The microneedles are hollow and/or porous to permit drug delivery at clinically relevant rates across skin or other tissue barriers, without damage, pain, or irritation to the tissue.
Other related microneedle devices in the patent literature are U.S. Pat. Nos. 5,527,288 and 5,611,806. More recently published patent applications on this topic are WO02/085446, WO02/085447, WO03/048031, WO03/053258 and WO02/100476A2.
These microneedles array patch technologies have achieved only limited success to date. Generally, there are a range of approaches configured to breach the stratum corneum to allow an enhanced take-up of drug in the viable epidermis. Although this has not been discussed in the patents referred to above, based upon reported research on ballistic particle delivery and cell death, the low strain rate of application, combined with the cases of smaller projections are likely to induce a lower incidence of cell death near the tips, than ballistic microparticle to delivery. Also, unlike ballistic microparticle delivery, these projections are removed from the tissue-alleviating the possibility of adverse effects of “carrier” materials delivered to the body, long term.
However, unlike biolistic targeting (FIG. 11(b)), and the direct injection of cells (FIG. 11(c)), these microneedle arrays do not have the advantage of readily and directly targeting inside the skin cells. This cellular/organelle targeting capability is key in a range of existing and potential methods of vaccination, gene therapy, cancer treatment and immunotherapy (Needle-free epidermal powder immunization. Chen et al, Expert Rev. Vaccines 1(3) p 265-276 (2002)) and diagnostic technologies.
Whilst U.S. Pat. No. 5,457,041 describes a patch for targeting cells, this is only suitable for use in vitro, and requires specialized apparatus to direct the micro-needles towards identified cells. The apparatus uses a microscope, to allow an operator to locate the cells in the sample tissue, and then direct the application of the micro-needles appropriately. As a result, this makes the device unsuitable for use in clinical environments, and limits the ability of the device to elicit a desired biological response.
Therefore, there still remains a need to provide projection-based technology which achieves a more accurately directed delivery of the active agent or stimulus to the desired site of action surrounding or within cells, without appreciable damage to them.